Electrodes with multilayer membranes and methods of using and making the electrodes

ABSTRACT

A sensor including a sensing layer is disposed over an electrode or an optode and a layer-by-layer assembled mass transport limiting membrane disposed over the sensing layer. The membrane includes at least one layer of a polyanionic or polycationic material. The assembled layers of the membrane are typically disposed in an alternating manner. The sensor also optionally includes a biocompatible membrane.

This application is a continuation of U.S. Ser. No. 09/854,310, filedMay 11, 2001, now U.S. Pat. No. 6,746,582, which claims the benefit ofU.S. Provisional Patent Application Ser. No. 60/203,762, filed May 12,2000. Each of these applications is incorporated herein by reference.

GOVERNMENT SUPPORT

This invention was made with government support under grant No.3R01DK42015 from the NIH-NIDDK. The government has certain rights in theinvention.

FIELD OF THE INVENTION

This invention relates to sensors and sensor components that havemultilayer membranes and methods of making and using the sensors andsensor components. In addition, the invention relates to enzymeelectrodes and optodes with multi-layer analyte-flux limiting membranesand methods of making and using the optodes and the electrodes.

BACKGROUND OF THE INVENTION

Miniature biosensors utilizing enzyme-containing optodes and electrodesfor monitoring biochemicals often include mass-transport controllingmembranes. The membranes can affect some or all of the characteristicsof the optodes or electrodes, including their sensitivity, size,apparent stability, dynamic range and selectivity. Micro-membranes foruse with miniature biosensors typically cannot be easily cut to sizefrom pre-formed membranes and if cut by a precision tool, such as alaser or an electron beam, their placement on and attachment to thesurface of an electrode or optode can be difficult.

The reproducible casting of micro-membranes can also be difficult. Forcast micro-membranes, the pore sizes and their distribution aretypically determined by the relative rates of nucleation andmass-transport during generation of the membrane by phase separation asa result of solvent evaporation. The outcome of the simultaneouslyoccurring nucleation and the mass transport processes depends on theevolution (meaning the time-dependence during the solvent evaporation)of the viscosity, the concentrations of the solvent and the non-solvent,and the membrane's leached phase and its residual phase. These areaffected by time-dependent temperature gradients and by thetime-dependent gradient of the partial pressure of the evaporatingsolvent over a droplet, the dimensions of which shrink and are afunction of the time-dependent contact angle with the wetted surface.

SUMMARY OF THE INVENTION

Generally, the present invention relates to electrodes and optodeshaving membranes to reduce analyte flux or reduce interferent flux orboth. One embodiment is a sensor that includes a sensing layer disposedon a substrate and a multilayer flux-limiting membrane disposed over thesensing layer. The membrane includes a first layer disposed on and boundto the sensing layer and one or more additional layers disposed on andbound to the preceding layers of the membrane. The substrate can have aconductive material upon which the sensing layer is disposed to form anelectrode or an optical material, such as an optical fiber, upon whichthe sensing layer is disposed to form an optode. As an example, themembrane includes at least two layers; one of which is a polycationiclayer or a polyanionic layer. Optionally, the membrane includes at leastone layer that has functional groups that can capture transition metalions.

Another embodiment is a method of making a sensor. A sensing layer isdisposed on a substrate. A first membrane layer is disposed on and bindsto the sensing layer. One or more subsequent membrane layers aredisposed over the first membrane layer, each of the subsequent membranelayers binding to the immediately preceding membrane layer. For example,the membrane layers can be formed by chemisorption or reactiveadsorption.

The above summary of the present invention is not intended to describeeach disclosed embodiment or every implementation of the presentinvention. The Figures and the detailed description which follow moreparticularly exemplify some but not all of these embodiments.

BRIEF DESCRIPTION OF THE DRAWINGS

The invention may be more completely understood in consideration of thefollowing detailed description of various embodiments of the inventionin connection with the accompanying drawings, in which:

FIG. 1 is a graph illustrating the dependence of the sensitivity of aglucose sensing electrode on the number of PAc/PAm/PAc/PAm/PAc/PVPEAsextets: (dashed line) none; (squares) one; (triangles) two; (circles)three;

FIG. 2 is a graph illustrating dependence of the apparent stability of aglucose sensing electrode on the number of PAc/PAm/PAc/PAm/PAc/PVPEAsextets: (a) none; (b) two; (c) three;

FIG. 3 is a graph illustrating current increments upon raising theconcentrations sequentially by 5 mM glucose, 5 mM glucose, 0.1 mMascorbate, 0.2 mM acetaminophen and 0.5 mM urate;

FIG. 4 is a graph illustrating changes in the current upon raising theglucose concentration in 5 mM increments, then adding Fe²⁺ in an amountthat in the absence of precipitation of iron phosphate would have raisedthe concentration of the cation to 0.1 mM with (a) ten PAc/PAm bilayersor (b) three PAc/PAm/PAc/PAm/PAc/PVPEA sextets;

FIG. 5 is a graph depicting in vivo experiments in which the glucoseconcentration was tracked by sensors with threePAc/PAm/PAc/PAm/PAc/PVPEA sextets. (a) (-) output of the sensorimplanted in the jugular vein; ( . . . ) output of the sensor implantedin the intrascapular subcutaneous tissue; (triangles) concentrations ofglucose in blood samples withdrawn from the contralateral jugular veinmeasured with a YSI glucose analyzer; (b) results of linear regressionanalysis of the data from the two sensors in FIG. 5( a): in the jugularvein (closed circles) and in the intrascapular subcutaneous tissue (opencircles);

FIG. 6 is a Clarke-type clinical error grid diagram of all data pointsassuming that the measured glucose concentration does not lag behind theblood glucose concentration at any time;

FIG. 7 is a Clarke-type clinical error grid diagram of all data pointsassuming that the sensor-measured glucose concentration lags by 3minutes behind the blood glucose concentration in the period of riseafter intravenous injection of a bolus of glucose and lags by 9 minutesduring the period of decline after intravenous injection of insulin; and

FIG. 8 is a schematic cross-sectional view of one embodiment of anelectrode, according to the invention.

While the invention is amenable to various modifications and alternativeforms, specifics thereof have been shown by way of example in thedrawings and will be described in detail. It should be understood,however, that the intention is not to limit the invention to theparticular embodiments described. On the contrary, the intention is tocover all modifications, equivalents, and alternatives falling withinthe spirit and scope of the invention.

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENT

The present invention is believed to be applicable to sensors,particularly miniature sensors, and methods of making and using thesensors. In particular, the present invention is directed to membranesfor electrodes and optodes for use in sensors and methods of making andusing the electrodes and optodes. While the present invention is not solimited, an appreciation of various aspects of the invention will begained through a discussion of the examples provided below.

A sensor includes an electrode or optode. Each electrode or optodeincludes a substrate. For an electrode, the substrate includes aconductive material that is typically either formed on a non-conductivesupport (e.g., a polymeric film) or as a wire, rod, plate, or otherobject. A sensing layer (e.g., transducing layer) is provided over atleast a portion of the conductive material to transduce the flux of achemical, usually a biochemical, termed the analyte to an electricalsignal, such as a current flowing through the electrode. An optode caninclude, for example, an optical fiber (or other optical substrate) atthe tip of which a layer of a biologically active macromolecule isprovided (preferably, immobilized). The sensing layer on the optodeconverts the change in the concentration of the analyte to a change inphoton flux.

The electrode or optode is covered with a membrane. The purpose of themembrane is to reduce analyte-flux or reduce or prevent interferent fluxto the electrode or a combination of these features. The membranecontains at least two layers, and, typically, at least six layers. Atleast one of the layers includes a layer containing a polyanion or apolycation and the polyanionic or polycationic layer is disposed overthe sensing layer. When operating to reduce flux of an analyte,interferent, or other material, the membrane typically provides a thinzone in which the solubility of a solute (e.g., analyte or interferent)is lower by at least an order of magnitude than it is in the samplesolution. A membrane having such a zone is referred to as “masstransport limiting”. The cross sectional area of the sensor throughwhich mass transport is limited is defined as the “active area”. Theactive area of the sensors in at least some embodiments of thisinvention is 0.1 cm² or smaller. For example, sensors can be formed withactive areas between 10⁻² cm² and 10⁻⁸ cm². Membranes having masstransporting areas (e.g., “active areas”) between 10⁻³ cm² and 10⁻⁵ cm²are particularly useful for biosensing applications.

The membrane can be formed, for example, by sequential chemisorption oflayers. Preferably, each particular layer, other than the terminal toplayer, binds both the preceding and succeeding layers of the membrane.Generally, two consecutive layers are not identical; however, layersmade of the same material but differently (for example, oppositely)oriented can be disposed next to each other. The layers typically forman array of bonds as a result of ionic, hydrophobic, coordinative,covalent, van der Waals, or hydrogen bonding interactions between thematerials of the two layers. For example, the membrane can be formed ofalternating polyanionic and polycationic layers.

In one embodiment, a micro-membrane is formed in-situ on a miniatureenzyme electrode by disposing (e.g., chemisorbing or otherwisedepositing) a polyanionic material on a polycationic surface of theelectrode, rinsing, disposing a polycationic on the polyanionicmaterial, rinsing, and repeating the cycle a desired number of times. Itwill be understood that a similar procedure can be used with apolyanionic surface of the electrode by first depositing a polycationicmaterial followed by a polyanionic material and repeating the cycle adesired number of times. Other orders of layers can also be used.

In particular, the membranes are useful in those biosensors thatfunction by chemically converting (e.g., reacting) a chemical orbiochemical. The chemical or biochemical can be, for example, an analytethat is being assayed, a product formed by reaction of the analyte, aco-reactant of the analyte, a product or reactant of a reaction that iscatalyzed or inhibited by the presence of the analyte, or a constituentwhose attachment to an optode or electrode is accelerated or inhibitedby the presence of the analyte.

As an example, membranes of the invention can be formed on, and areevaluated in the Examples below for, miniature glucose oxidizing orelectrooxidizing electrodes. For these electrodes, glucose is typicallyconverted to gluconolactone in the first step of a detection reaction.Suitable miniature glucose electrodes include, for example, thosedisclosed in U.S. Pat. Nos. 5,262,305; 5,262,035; 5,264,104; 5,593,852;5,665,222; 6,143,164; 6,120,676; 6,576,101; 6,134,461; 6,338,790, andU.S. patent application Ser. No. 09/434,026, all of which areincorporated herein by reference. It will be appreciated that electrodesfor the detection of analytes other than glucose can also benefit fromthe use of the membranes described herein. Glucose electrodes areillustrated herein as an application example.

Glycemia has been monitored amperometrically in the subcutaneousinterstitial fluid with miniature electrodes for some time. Theelectrode reactions applied in such monitoring include (a) the mediatedelectrooxidation of glucose to gluconolactone at electrodes coated witha redox mediator (e.g., a redox polymer which electrically “wires” thereaction centers of glucose oxidase to an electrode) (Equations 1a and1b, below) or, alternatively, (b) the glucose oxidase catalyzed reactionof glucose with O₂, to produce gluconolactone and H₂O₂ (Equation 2a,below), followed either by electro-oxidation of the H₂O₂ (Equation 2b,below), or by monitoring the change in the O₂ partial pressure orconcentration.glucose+2 bound mediator(ox)→gluconolactone+2 boundmediator(red)+2H⁺  (1a)2 bound mediator (red)→2 bound mediator (ox)+2e ⁻  (1b)glucose+O₂→gluconolactone+H₂O₂  (2a)H₂O₂ O→O₂+2H⁺+2e ⁻  (2b)The current at the electrode generally scales with the flux of glucoseto the electrode. The flux, and therefore also the current, typicallyincreases linearly with the glucose concentration as long as the entireglucose flux at the electrode is consumed in the electrode reaction.

FIG. 8 illustrates one example of a suitable electrode 100. Theelectrode includes a conductive material 102 within an insulating sleeve104. Disposed on the conductive material 102 are a sensing layer 106, amembrane 108, and a biocompatible layer 110. Examples of similarelectrode configurations, together with examples of some suitablesensing layers and biocompatible layers are provided in U.S. Pat. No.5,593,852. Other suitable sensing layers and biocompatible layers aredescribed, for example, in U.S. Pat. Nos. 5,665,222; 6,143,164;6,120,676; 6,576,101; 6,134,461; 6,338,790, and U.S. patent applicationSer. No. 09/434,026, all of which are incorporated by reference.

The conductive material 102 can be, for example, carbon, a metal, or aconductive compound or polymer. The insulating sleeve 104 is typicallyformed from an insulating compound or polymer. In one method ofmanufacture, a tip of a metal wire (e.g., a gold wire) is etched to forma recess within the insulating sleeve where the sensing layer 106,membrane 108, and biocompatible layer 110 are deposited.

The sensing layer 106 typically provides a mechanism for transducing ananalyte flux (or the flux of a product molecule formed or consumed by areaction of the analyte) to an electrical signal. The sensing layer cancontain, for example, a redox mediator to facilitate the indirect ordirect transfer of electrons between the conductive material and theanalyte. One type of redox mediator is a transition metal complex orcompound (e.g., an osmium, ruthenium, or iron complex or compound). Theredox mediator can be a monomeric redox compound or complex, but ispreferably in a non-leachable form, such as a redox polymer. The redoxpolymer, has a polymeric backbone with multiple redox centers. Examplesof suitable redox mediators are disclosed in, for example, U.S. Pat.Nos. 5,262,305; 5,262,035; 5,264,104; 5,593,852; 5,665,222; 6,143,164;6,120,676; 6,576,101; 6,134,461; 6,338,790, and U.S. patent applicationSer. No. 09/434,026, all of which are incorporated by reference.

The sensing layer 106 can also include a second electron transfer agent,such as an enzyme. The second electron transfer agent can catalyze theelectrochemical oxidation or reduction of the analyte. As an example,suitable second electron transfer agents for glucose include glucoseoxidase or glucose dehydrogenase; for lactate, lactate dehydrogenase;and for hydrogen peroxide, peroxidase.

The biocompatible layer 110 prevents the penetration of largebiomolecules into the electrodes. This can be accomplished by using abiocompatible layer 110 having a pore size that is smaller than thebiomolecules that are to be excluded. Such biomolecules can foul theconductive material 102 or the sensing layer 106 thereby reducing theeffectiveness of the electrode 100 and altering the expected signalamplitude for a given analyte concentration. The biocompatible layer 110may also prevent protein adhesion to the electrode 100, formation ofblood clots, and other undesirable interactions between the electrode100 and body. A preferred biocompatible coating is a hydrogel, whichcontains at least 20 wt. % fluid when in equilibrium with theanalyte-containing fluid. Examples of suitable hydrogels are describedin U.S. Pat. No. 5,593,852, incorporated herein by reference, andinclude crosslinked polyethylene oxides, such as polyethylene oxidetetraacrylate.

The preferred membrane 108 includes at least three, and typically atleast six, twelve, or eighteen layers. One example of a membrane 108 isformed using alternating polycationic and polyanionic layers. Typically,these layers are formed using polymers. Suitable polycationic polymersinclude, for example, polyallylamine hydrochloride (PAm),poly(4-vinylpyridine) quaternized by reacting about one third to onetenth of the pyridine nitrogens with 2-bromoethylamine (PVPEA),polyethylene imine, and polystyrene modified with quaternary ammoniumfunctions. Suitable polyanionic polymers include, for example,poly(acrylic acid) (PAc), poly(methacrylic acid), partially sulfonatedpolystyrene, polystyrene modified with functions having carboxylateanions, and DNA (deoxyribonucleic acid) or RNA (ribonucleic acid)strands, fragments or oligomers. The membrane 108 can serve one or morefunctions including, for example, a) limiting glucose flux or b)reducing or eliminating the flux of interferents to the electrode.Glucose sensors will be described as an example, but sensors can beformed for other analytes using the same principles.

In glucose sensing optodes or electrodes, a glucose flux-limitingmembrane can enhance one or more properties of the optode or theelectrode including, for example, expanding the dynamic range, enhancingthe apparent stability and improving the selectivity for glucose whichmay enable one-point calibration of an implanted electrode.

Reducing the flux of glucose (or any other analyte) using a membrane canexpand the dynamic range of a sensor. The upper limit of the dynamicrange of the sensor is that glucose concentration where the entireinflux of the analyte is still consumed. When this limit is exceeded,the rate of glucose-conversion is slower than the influx and the sensor“saturates”: the current no longer increases when the glucoseconcentration is raised. The kinetic limit, represented by the currentdensity at the glucose concentration above which it no longer increases,is an intrinsic property of the electrode. This current densitytypically scales linearly with the rate of the slower of reactions 1(a)or 1(b) for a mediator-comprising glucose oxidase electrode or theslower of reactions 2(a) or 2(b) in an O₂— utilizing electrode. Becausethe membrane does not affect the intrinsic rate of the slowest step, itsinsertion between the assayed fluid and the electrode expands the upperlimit of the dynamic range.

Typically, the sensor's apparent stability, which is the stabilityperceived by its user, can also be improved upon insertion of theglucose flux limiting membrane. The user remains unaware of thedeterioration of the electrode's chemistry as long as the kinetics ofthe slowest step remains fast enough to convert all the influx ofglucose to a current. Furthermore, the more blocking the membrane is,the better the apparent stability will typically be. The increasedapparent stability and the upward extension of the dynamic range aregained at the cost of reduced specific sensitivity, defined as thecurrent per unit area at unit glucose concentration. The specificsensitivity decreases because less glucose reaches the reactive zone onthe electrode. When the specific sensitivity decreases, a largerglucose-transporting area may be needed for the sensor's current toreach an easily measured value.

The membrane thus defines, at least in part, (a) the lowest analyteconcentration where the current is large enough to be easily measuredand the highest analyte concentration above which the electrode'scurrent no longer increases with the concentration of the analyte; (b)the extent of the feasible miniaturization of the sensor; and (c) theapparent stability of the sensor. Fitting of these characteristics tothose sought in the intended application requires tailoring of theflux-controlling membrane. While the characteristics of the earliersolvent-cast membranes depended on the spreading of the castingsolution, and on the gradients of temperature and of solvent partialpressure near the membrane being formed; the characteristics of themembranes formed according to the invention typically do not depend onthese factors. As a result, the simultaneous tailoring of their manydesired characteristics is typically easier.

In at least some embodiments, the membrane also reduces or eliminatesthe flux of interferents to the conductive material of the electrode.The bias resulting from the presence of an interferent that is rapidlyelectrooxidized at the applied potential is defined by its flux to theelectrode surface. In the absence of a membrane, the flux is controlledby diffusion in the solution and is defined by the diffusivity and theconcentration of the interferent. When a membrane is inserted betweenthe solution and the electrode, the % bias is defined by the ratio ofthe products of the solution concentrations and permeabilities of theinterferent and glucose.

$\text{\%~~bias} = {\frac{C_{I,S} \cdot D_{I,S}}{C_{G,S} \cdot D_{G,S}} \times 100\%\mspace{14mu}\text{(solution)}}$$\text{\%~~bias} = {\frac{C_{I,S} \cdot P_{I,M}}{C_{G,S} \cdot P_{G,M}} \times 100\%\mspace{14mu}\text{(membrane)}}$$P_{I,M} = {{K_{I,M}{D_{I,M}/l_{M}}} = {\frac{C_{I,M}}{C_{I,S}} \times {D_{I,M}/l_{M}}}}$where C is the concentration of the interferent (I) or glucose (G) inthe solution (S) or membrane (M), D represents the appropriatediffusivity, P represents the partition coefficient, and 1_(M)represents the thickness of the membrane. Because the permeability is aproduct of the concentration of the diffusing species in the membraneand of its diffusivity in the membrane, the % bias increases linearlywith the ratio of the partition coefficients of the interferent and ofglucose between the membrane and the solution. When only monovalentanionic species (Cl⁻, ascorbate, urate) are present, the concentrationof anions in the polycationic POs-EA membrane of the sensing layerequals the concentration of its cationic charges, which is about anorder of magnitude higher than the concentration of anions in thesolution. As a result, the sensing layer is permselective for anionsover glucose and in the absence of a neutral or polyanionicglucose-flux-controlling membrane (e.g. Nafion™) the flux andelectro-oxidation current of 0.1 mM ascorbate could equal or exceed 10%of that of 10 mM glucose. The membranes described herein are nottypically permselective for anions. As a result, ascorbate or urate isnot preferentially electrooxidized. As seen in FIG. 3, at 10 mM glucoseconcentration the combined % bias resulting of the presence of 0.1 mMascorbate, 0.2 mM acetaminophen and 0.5 mM urate is less than 5%, eventhough the sensor is poised at 450 mV (Ag/AgCl), a potential whereacetaminophen is not rapidly electrooxidized, but ascorbate and urateare.

Transition metal ions also influence the sensor readings. Transitionmetal ions reduce the intrinsic kinetic capacity of the enzyme layer toelectrooxidize glucose and thereby severely reduce the dynamic range andthe sensitivity. (See FIG. 4) The loss is attributed to both inhibitionof the enzyme and a reduction in the surface density of electroactiveredox centers caused by excessive crosslinking through coordination ofpyridine rings of neighboring redox centers. In the Examples, atransition metal ion capturing PVPEA/PAc bilayer is included in themembrane. The addition of the PVPEA layer provides an alternative sitefor capture of transition metal ions by providing pyridinefunctionalities that can complex with the transition metal ions. It willbe recognized that other materials can be used in place of PVPEA.Typically, the replacement materials will include functionalities thatcan form complexes with the transition metal ions. Because the PVPEAlayer is already highly crosslinked, the incremental crosslinking bycoordination of the transition metal ions does not change excessivelythe permeability of the micro-membrane to glucose.

As an example, described in the Examples section below, membranes wereassembled by sequentially chemisorbing polyanionic and polycationicmaterials on miniature (5×10⁻⁴ cm²) enzyme electrodes. The sequentialchemisorption process allowed the simultaneous tailoring of theirsensitivity, dynamic range, drift and selectivity. When assembled ontips of 250 μm diameter gold wires coated with a redox polymer/glucoseoxidase sensing layer, they allowed tailoring of the glucose electrodesfor greater than 2 nA/mM sensitivity; 0 to 30 mM dynamic range; drift of≦5% per 24 hours at 37° C. at 15 mM glucose concentration; ≦5% currentincrement by the combination of 0.1 mM ascorbate, 0.2 mM acetaminophenand 0.5 mM urate. The membranes also retained transition metal ions thatotherwise bind to and damage the redox polymer and the enzyme. Theelectrodes were tested in the jugular veins and in the intrascapularsubcutaneous region of anaesthetized and heparinized non-diabeticSprague-Dawley rats, in which rapid changes of glycemia were forced byintravenous glucose and insulin. After one-point in-vivo calibration ofthe electrodes, all of 152 data points were clinically accurate when itwas assumed that after insulin injection the glycemia in thesubcutaneous fluid lags by 9 minutes behind that of blood withdrawn fromthe insulin-injected vein.

EXAMPLES

The following are the materials used to make the electrodes: Glucoseoxidase (GOx) (Fluka, Milwaukee, Wis., EC 1.1.3.4, 197 units/mg) fromAspergillus niger; poly (ethylene glycol) diglycidyl ether (400)(PEGDGE), and polyallylamine hydrochloride (MW 50,000) (PAm) fromPolysciences, Warrington, Pa.; and polyacrylic acid sodium salt (PAc)(MW 15,000), 1-[3-(dimethylamino)propyl]-3-ethylcarbodiimide (EDC), andN-hydroxysuccinimide (NHS) from Aldrich, Milwaukee, Wis. The redoxpolymer, poly (4-vinylpyridine) partially N-complexed with[Os(bpy)₂Cl]^(+/2+) and quaternized with bromoethylamine (POs-EA), wasprepared as described in Gregg et al., J. Phys. Chem. 95, 5970 (1991),incorporated herein by reference. The related polymer from which theosmium complex was omitted, PVPEA, was prepared by quaternizingpoly(4-vinylpyridine) with bromoethylamine.

Example 1

Formation of Glucose Electrodes. Miniature gold electrodes werestructurally similar to those described in Csoregi et al., Anal. Chem.66, 3131 (1994) and U.S. Pat. No. 5,593,852, both of which areincorporated herein by reference. The electrodes were made ofpolyimide-insulated 0.25 mm gold wire, which was cut to 5 cm longpieces. At one end, the insulation was stripped from 0.5 cm of the wireto make the electrical contact. At the other end, a 90 μm deeppolyimide-walled recess was formed by electrochemically etching away thegold under galvanostatic control by an EG&G PARC 273Apotentiostat/galvanostat.

The tip of the gold wire at the bottom of the shielded recess was coatedwith the transduction (sensing) layer; a micro-membrane; and abiocompatible layer. The first and third layers were formed bymicropipetting polymer solutions onto the gold surface under amicroscope, using a micromanipulator. The micro-membrane was formed bydip and rinse cycles.

The sensing layer included the redox polymer POs-EA and GOx crosslinkedwith PEGDGE. A 20 mg/mL solution of GOx were dissolved in a 0.1 M sodiumbicarbonate aqueous solution. The GOx solution was then mixed at 2:1volume ratio with a 12 mg/mL solution of sodium periodate and themixture was reacted in the dark at room temperature for 1 hour. 2 μL ofthe now periodate-oxidized GOx solution was mixed with 16 μL of 10 mg/mLPOs-EA solution and 1.4 μL of 2.5 mg/mL PEGDGE solution. 15 droplets ofabout 5 nL mixed solution were sequentially micropipetted into therecessed cavity formed by back-etching the gold in its polyimideinsulation. The resulting films were cured at 45° C. for 30 minutes.

The micro-membrane was formed over the sensing layer. Thepolyelectrolyte solutions 20 mM PAc, 20 mM PAm and 20 mM PVPEA (theconcentrations being those of the acidic or the basic functions, not ofthe macromolecules) were prepared in 0.1 M NaCl buffered at pH 6 with0.1 M 2-[N-morpholino] ethanesulfonic acid (MES). This buffer was usedalso to prepare 20 mM EDC and 50 mM NHS. The sensors were coated bydipping and rinsing cycles, alternately in PAc and in PAm, to formPAc/PAm bilayers, or in PAc then in PVPEA to form the PAc/PVPEAbilayers. The sequence of the resulting sextets wasPAc/PAm/PAc/PAm/PAc/PVPEA; the slashes (/) representing rinses with MESbuffer to remove the excess (unbound) polyelectrolyte. All of thesensors used in vitro and in animals, except for those made for theparametric optimization of the sensors, had three of the sextet layers(18 layers total).

The biocompatible layer was formed over the micro-membrane. Thebiocompatible layer was formed by UV-photocrosslinking tetraacrylatedPEO, using 2,2-dimethoxy-2-phenyl-acetophenone as the photoinitiator.

Example 2

In Vitro Experiments using the Electrode of Example 1. In vitroexperiments were carried out in a stirred, water-jacketedelectrochemical cell in 0.15 M NaCl, 0.02 M phosphate buffer solutionwith pH 7.1. The cell had a saturated Ag/AgCl reference electrode, aplatinum counter electrode and the modified 0.25 mm gold wire tipworking electrode, as described in Example 1. Unless otherwise statedthe working electrode was poised at 400 mV vs. Ag/AgCl, and the cell wasmaintained at 37° C. with an isothermal circulator (Fisher Scientific,Pittsburgh, Pa.). The potential was controlled by a CHI832electrochemical detector (CH Instrument, Austin, Tex.) and a PCcollected the data.

FIG. 1 illustrates the dependence of the sensitivity on the number ofPAc/PAm/PAc/PAm/PAc/PVPEA sextets where the dashed line indicates nomicro-membrane, the squares indicate one sextet, the triangles indicatetwo sextets, and the circles indicate three sextets. This demonstratesthe expansion of the linear range and the corresponding decrease insensitivity when an increasing number of countercharged polyelectrolytelayers is applied on the sensing layer. When the micro-membraneconsisted of two sextets (of the sequencePAc/PAm/PAc/PAm/PAc/PVPEA/PAc/PAm/PAc/PAm/PAc/PVPEA) the currentincreased linearly with the glucose concentration at least up to 30 mM.As seen in FIG. 1, the linear domain in-vitro now extended through theentire physiologically relevant 2-30 mM range.

FIG. 2 illustrates the dependence of the apparent stability on thenumber of PAc/PAm/PAc/PAm/PAc/PVPEA sextets where line (a) indicates nomicro-membrane, line (b) indicates two sextets, and line (c) indicatesthree sextets. FIG. 2 illustrates the improvement in the stability ofthe current at 15 mM glucose concentration and at 37° C. when the numberof the layers was increased. In the absence of a micro-membrane, 39% ofthe current was lost in the initial 24 hour period. Application of twosextets (PAc/PAm/PAc/PAm/PAc/PVPEA/PAc/PAm/PAc/PAm/PAc/PVPEA) reducedthe loss to 9%. When three sextets were applied(PAc/PAm/PAc/PAm/PAc/PVPEA/PAc/PAm/PAc/PAm/PAc/PVPEA/PAc/PAm/PAc/PAm/PAc/PVPEA),the 24 hour loss dropped to 5%.

Both urate and the ascorbate anions are electrooxidized at potentialspositive of 200 mV (SCE). In the absence of a micro-membrane theelectrooxidation current of anionic interferents is disproportionatelyhigh when the redox-polymer backbone is a polycation. For example, theascorbate electrooxidation current at 0.1 mM ascorbate concentration isgreater than the glucose electrooxidation current at 1 mM concentration.The cause of the disproportionate electrooxidation of anionicinterferents is thought to be due to the scaling of their concentrationwithin the redox polymer with the density of cationic sites. As aresult, the permeability of the membrane to anionic interferents (whichis the product of concentration and diffusivity) is higher than that ofneutral molecules like glucose.

Application of the micro-membrane alleviated the disproportionatelylarge interference by ascorbate and urate, as shown in FIG. 3. At 10 mMglucose concentration, the aggregate increase in current produced by thecombination of 0.1 mM ascorbate, 0.2 mM acetaminophen and 0.5 mM uratewas less than 5%. The sensitivity of the electrodes to glucose was notchanged.

Transition metal ions like Zn²⁺ and Fe^(2+/3+), which are present inserum at <50 μM concentration, can reduce the sensitivity of the “wired”enzyme electrode. The loss is observed even in the presence of 20 mMphosphate, which precipitates most of these ions at neutral pH.Application of multiple bilayers consisting only of PAc/PAm, did notappear to alleviate this loss. The loss was, however, alleviated byincorporating PAc/PVPEA bilayers, which had transition-ion complexingpyridine functions as shown in FIG. 4.

The results show that the following sensor characteristics can besimultaneously provided by a tailored micro-membrane: Linear range, 2 to30 mM; sensitivity per unit area of 4 to 6 μAcm⁻²mM⁻¹, translating to 2to 3 nA mM⁻¹ sensitivity for an electrode having a 5×10⁻⁴ cm² masstransporting and sensing area; and ≦5% loss in sensitivity in 24 hrs.

Example 3

In Vivo Experiments using the Electrode of Example 1. MaleSprague-Dawley rats, 400-500 g, were pre-anesthetized with halothane(Halocarbon Laboratories, North Augusta, S.C.) and anesthetized byintraperitoneal injection (0.5 mL) of a solution made of equal volumesof acepromazine maleate (10 mg/mL), ketamine (100 mg/mL), and xylazine(20 mg/mL). The animals were shaved about their necks, abdomens, andbetween their scapulae, and then secured on a homeothermic blanketsystem (Harvard Apparatus, South Natick, Mass.). A 0.0375-in.-diametermedical grade silicone tube was inserted into the proximal portion oftheir right external jugular vein and secured with 4-0 silk sutures. Adose of 100 units/kg body weight of heparin solution was thenadministered, followed by an equal volume of saline to clear the line. Aglucose sensor was implanted subcutaneously between the scapulae, usinga 22-gauge Per-Q-Cath introducer (Gesco International, San Antonio,Tex.). The sensor was taped to the skin to prevent its movement. Asecond silicone tubing of ˜2 cm length was inserted into the proximalside of the left external jugular vein as a guide, and the secondglucose sensor was inserted inside the guide tube. The tube and thesensor were then secured with a microvascular clamp, with the sensorprotruding beyond the end of the guide tube. A Ag/AgCl surface skinreference electrode was attached to the animal's abdomen afterconductive gel was applied. The sensors and the reference electrode werethen connected to an EG&G PARC Model 400 bipotentiostat, the output ofwhich was recorded with a Rustrak Ranger data-logger (Rustrak Ranger,East Greenwich, R.I.). Data collection started 40-60 minutes after thesensors were poised at +450 mV vs. Ag/AgCl. The reference blood sampleswere collected from the right jugular vein and were analyzed using a YSIModel 2300 glucose analyzer (YSI, Yellow Spring, Ohio).

In each experiment a 50% glucose solution (300 mg/kg) was administeredintravenously to induce a rapid rise in glucose concentration. A rapiddecline in glucose concentration was then induced by an intravenousinsulin injection (regular U-100, 0.5 unit/kg). At the end of theexperiment, the rat was euthanized by intravenous sodium pentobarbitalinjection, consistent with the recommendations of the panel onEuthanasia of the American Veterinary Association. The protocols of theexperiments in vivo were approved by the University of TexasInstitutional Animal Use and Care Committee.

The experiment was started when the glycemia of the rats was at a steadystate, the steady glucose concentration being confirmed by withdrawingthree blood samples and their analysis with the YSI analyzer. Twominutes after the third withdrawal a bolus of glucose was injected. Thesensors were calibrated in vivo by independently analyzing a singlesample of blood about 2 minutes before the injection of the bolus ofglucose. FIG. 5( a) shows the variation in the sensor-measured glycemiaafter boli of glucose and insulin were sequentially administeredintravenously. FIG. 5( b) shows the correlation of the sensor readingsand the YSI results of FIG. 5( a). The linear regressions for the eightsensors are summarized in Table 1, below. The average correlationcoefficient (r²) was 0.960 for the jugular vein sensors and 0.935 forthe subcutaneous sensors. The average of the intercept was 0.6 mM±0.6mM, not differing greatly from the reported −0.79 mM to +0.48 mM rangeof intercepts of home blood glucose meters used by self-monitoringdiabetic patients.

As seen in FIG. 5 and in Table 1 the subcutaneous and the jugular-veinimplanted sensors with the in situ synthesized micro-membranesaccurately track the YSI-glucose analyzer measured blood glucoseconcentration when calibrated in vivo at one point. The clinicalvalidity of glucose assays is often judged by their position in zones ofthe Clarke plot. Points in zone A of the Clarke plot represent accurateassays. Points in zone B represent less accurate assays leading to validclinical decision. Points in Zone C reflect assays leading toinappropriate, though not harmful, clinical decisions. Points in zone Dreflect assays leading to the missing of a necessary clinical action(consumption of a glucose-source or insulin-injection) when such actionis required. Points in Zone E reflect assays leading to clinical actionthat are the opposite of the required, such as assays indicating theneed to inject insulin when the patient is already hypoglycemic.

In absence of correction for the potential transient difference betweenthe blood and the subcutaneous glucose concentrations during rapid riseor decline periods the Clarke-type error grid analysis (FIG. 6) of thedata shows that 95.5% of the points are in the clinically accurate oracceptable zones A or B (Table 2, below). The points in zone D (4.5%)resulted of failure to detect hypoglycemia and originated in periods ofrapid decline following intravenous administration of insulin. Thefraction of points in zone D was reduced to less than 1% when it wasassumed that following insulin injection, but not after glucoseinjection, the subcutaneous glucose concentration lags behind that ofthe insulin-injected vein by 9 minutes. Table 3, below, shows that thefraction of points in zones other than zone A of the Clarke plotincreased when it was assumed that in the period after intravenousadministration of glucose the lag time of the subcutaneous glucoseconcentration behind that in the glucose-injected vein was greater than0 to 3 minutes; when no lag or a 3 min lag were assumed all points werein zone A.

The assumptions of a 0 to 3 min lag of the subcutaneous glycemia afterglucose injection and of a 9 min lag after insulin injection not onlybrought all of the 88 points measured with the subcutaneous electrodeinto zones A and B of the Clarke plot, but also increased the ratio ofzone A to zone B points. For the intravenous sensor-measured glycemia,the assumption of a 3 min lag of the contralateral venous glycemiabehind that in the injected vein, whether after glucose or insulininjection, brought all points into zones A and B of the Clarke plot(Tables 3 and 4, below). The assumptions of a 9 min lag in thesubcutaneous glycemia after insulin injection and of a 3 min lag in thesubcutaneous glycemia after glucose injection, as well as in thecontralateral venous glycemia after glucose or insulin injection, thusbrought 163 of the 176 points (92.6%) into zone A; 13 points being inzone B. Comparison of the values in Table 5, below, with those in Table1 summarizes the effect of these assumptions on the distribution ofpoints, the slopes, the intercepts, and the percent difference betweenthe YSI and the sensor readings. The sensors with the in-situ assembledmicro-membranes accurately measured the glycemia in the jugular vein andin the interstitial subcutaneous fluid.

The present invention should not be considered limited to the particularexamples described above, but rather should be understood to cover allaspects of the invention as fairly set out in the attached claims.Various modifications, equivalent processes, as well as numerousstructures to which the present invention may be applicable will bereadily apparent to those of skill in the art to which the presentinvention is directed upon review of the instant specification.

TABLE 1 Results of linear regression analysis of the correlation betweenthe actual glucose blood glucose concentrations and concentrationsmeasured by the implanted sensors*^(a) Intercept % Slope (mg/dL) r²difference*^(b) jugular vein 0.83  9.7 0.966 −3.3 jugular vein 0.96 15.00.944 11.5 jugular vein 1.02 −4.2 0.957 −2.7 jugular vein 0.94 −0.90.972 −9.2 subcutaneous 0.85 13.1 0.943  3.3 subcutaneous 0.90 30.70.938 12.7 subcutaneous 0.90  3.5 0.928 −9.0 subcutaneous 0.86 15.70.931  0.1 Average 0.91 ± 0.06 10.3 ± 11.1 0.947 ± 0.016  0.4 ± 8.3*^(a)22 blood samples were withdrawn and independently analyzed in eachexperiment. *^(b)% difference = [Σ((sensor glucose-blood glucose)/bloodglucose)]/n

TABLE 2 Clarke-type error grid analysis of all data, without and withassumption of lag. No lag assumed Lag assumed Zone Data points % Datapoints % A 151 85.8 163 92.6 B 17 9.7 13 7.4 C 0 0.0 0 0.0 D 8 4.5 0 0.0E 0 0.0 0 0.0 *A 3 min lag was assumed, except for the subcutaneoussensors after insulin injection, for which a 9 min lag was assumed.

TABLE 3 Dependence of the distribution of the data points between thezones of the Clarke plot on the assumed lag after injection of glucose.Time Delay (mins) 0 3 5 7 9 11 Intravenous Sensors Zone A 32 31 30 30 2624 Zone B 0 1 2 2 6 7 Zone D 0 0 0 0 0 1 Subcutaneous Sensors Zone A 3232 30 30 29 26 Zone B 0 0 2 2 3 6 Zone D 0 0 0 0 0 0

TABLE 4 Dependence of the distribution of the data points between thezones of the Clarke plot on the assumed lag after injection of insulin.Time Delay (mins) 0 3 5 7 9 11 Intravenous Sensors Zone A 35 41 40 38 3433 Zone B 7 3 4 6 10 11 Zone D 2 0 0 0 0 0 Subcutaneous Sensors Zone A28 34 36 35 37 34 Zone B 10 7 6 7 7 10 Zone D 6 3 2 2 0 0

TABLE 5 Results of linear regression analyses of the correlation betweenthe YSI-measured venous glycemia and the sensor-measured glycemiaassuming the optimal lag times.* Slope Intercept, mg/dL r² % Differencejugular vein 0.78 9.8 0.964 12.5 jugular vein 0.99 7.3 0.968 9.9 jugularvein 1.03 −11.6 0.981 5.7 jugular vein 0.94 −5.8 0.980 9.6 average 0.94± 0.07 2.4 ± 8.6 0.973 ± 9.4 ± 1.9 0.007 subcutaneous 0.85 7.5 0.94710.7 subcutaneous 0.97 −23.8 0.954 16.9 subcutaneous 0.97 3.8 0.975 5.7subcutaneous 0.88 2.3 0.973 9.0 average 0.92 ± 0.05 −2.5 ± 7.2 0.962 ±10.6 ± 3.0 0.012 average 0.93 ± 0.08  −1.3 ± 11.6 0.968 ± 10.0 ± 3.6(all data) 0.012 *3 min lag for the contra-lateral venous and thesubcutaneous glycemia after injection of glucose; 3 min lag of thecontra-lateral venous glycemia after injection of insulin; 9 min lag ofthe subcutaneous glycemia after injection of insulin.

1. A sensor, comprising: a sensing layer disposed on a substrate; amultilayer flux-limiting membrane disposed on the sensing layer, themembrane comprising a sextext layer comprising six alternatingpolycationic and polyanionic layers, each layer disposed on and bound tothe immediately preceding layer of the membrane, wherein at least one ofthe polyanionic layers and the polycationic layers comprises pyridine.2. The sensor of claim 1, wherein the sensing layer comprising an osmiumredox polymer.
 3. The sensor of claim 1, wherein the sensing layercomprises a non-leachable redox polymer.
 4. The sensor of claim 1,wherein the sensing layer comprises an enzyme.
 5. The sensor of claim 1,wherein the flux-limiting membrane is disposed on and bound to thesensing layer.
 6. The sensor of claim 1, further comprising abiocompatible layer disposed over the multilayer membrane.
 7. The sensorof claim 1, wherein the multilayer flux-limiting membrane comprises asecond-sextet layer.
 8. The sensor of claim 1, wherein the substratecomprises an electrode upon which the sensing layer is disposed.
 9. Thesensor of claim 1, wherein the substrate comprises an optode upon whichthe sensing layer is disposed.
 10. The sensor of claim 1, wherein thesensor has an active area of 0.1 cm² or less.
 11. The sensor of claim10, wherein the active area is 10⁻² cm² to 10⁻⁸ cm².
 12. A sensor,comprising: an osmium sensing layer disposed on a substrate; amultilayer flux-limiting membrane disposed on the sensing layer, themembrane comprising a first polyanionic layer and a first polycationiclayer, the first polyanionic layer disposed on the sensing layer,wherein at least one of the first polyanionic layer and a firstpolycationic layer comprising pyridine; and a biocompatible layerdisposed over the multilayer membrane.
 13. The sensor of claim 12,wherein each of the first polyanionic layer and a first polycationiclayer comprising pyridine.
 14. The sensor of claim 12, wherein themultilayer flux-limiting member further comprises a second polyanioniclayer disposed on the first polycationic layer, and the biocompatiblelayer being disclosed on the second polyanionic layer.
 15. The sensor ofclaim 12, wherein the multilayer flux-limiting member further comprisesa second polyanionic layer and a second polycationic layer, a secondpolyanionic layer disposed on the first polycationic layer, and thebiocompatible layer being disposed on the second polycationic layer. 16.A sensor, comprising: an osmium sensing layer disposed on a substrate; amultilayer flux-limiting membrane disposed on the sensing layer, themembrane comprising a first polycationic layer and a first polyanioniclayer, the first polycationic layer disposed on the sensing layer, atleast one of the first polycationic layer and a first polyanionic layercomprising pyridine; and a biocompatible layer disposed over themultilayer membrane.
 17. The sensor of claim 16, wherein each of thefirst polyanionic layer and a first polycationic layer comprisespyridine.
 18. The sensor of claim 16, wherein the multilayerflux-limiting member further comprises a second polycationic layerdisposed on the first polyanionic layer, and the biocompatible layerbeing disposed on the second polycationic layer.
 19. The sensor of claim16, wherein the multilayer flux-limiting member further comprises asecond polycationic layer and a second polyanionic layer, a secondpolycationic layer disposed on the first polyanionic layer, and thebiocompatible layer being disposed on the second polyanionic layer. 20.A sensor, comprising: a sensing layer disposed on a substrate; amultilayer flux-limiting membrane disposed on the sensing layer, themembrane comprises alternating polycationic and polyanionic layers, eachlayer disposed on and bound to the immediately preceding layer of themembrane, wherein at least one of the polyanionic layers and thepolycationic layers comprises pyridine.
 21. The sensor of claim 20,wherein the sensing layer comprising an osmium redox polymer.
 22. Thesensor of claim 20, wherein the sensing layer comprises a non-leachableredox polymer.
 23. The sensor of claim 20, wherein the sensing layercomprises an enzyme.
 24. The sensor of claim 20, wherein theflux-limiting membrane is disposed on and bound to the sensing layer.25. The sensor of claim 20, wherein at least one of the polyanioniclayers and the polycationic layers comprises pyridine.
 26. The sensor ofclaim 20, further comprising a biocompatible layer disposed over themultilayer membrane.
 27. The sensor of claim 20, wherein the substratecomprises an electrode upon which the sensing layer is disposed.
 28. Thesensor of claim 20, wherein the substrate comprises an optode upon whichthe sensing layer is disposed.
 29. The sensor of claim 20, wherein thesensor has an active area of 0.1 cm² or less.
 30. The sensor of claim29, wherein the active area is 10² cm² to 10⁻⁸ cm².